The present invention relates to a magnetic resonance imaging (MRI) system which causes an object under examination to undergo a magnetic resonance (MR) phenomenon, detects an MR signal based on the MR phenomenon, and performs processing including image reconstruction for the detected data, thereby obtaining MR image data representing the distribution of MR data in a selected slice of the object, e.g., the density of a specific nuclear spin and/or relaxation time constant, and the like. More particularly, the invention relates to a method and system for MRI for performing a predetermined arithmetic operation with the plurality of MR images to calculate another MR image.
In an MRI system, an object to be examined is placed in a uniform static magnetic field, a gradient field is superposed on the uniform magnetic field, and an excitation rotating field is applied to the object in the magnetic field, thus causing an MR phenomenon in a prospective slice portion of the object. An MR signal caused by the MR phenomenon is detected, and is subjected to predetermined processing including image reconstruction, thus obtaining slice data which reflects MR data in the slice of the object. The MRI system is widely used as a medical diagnosis system.
An imaging operation in a conventional MRI system will now be briefly described with reference to FIG. 1.
Object P to be examined (i.e., a patient) is placed in a sufficiently uniform static magnetic field H0, along a z axis in FIG. 1. Static magnetic field H0 is generated by a magnetic field generator (not shown). Linear magnetic gradient Gz is applied to static magnetic field H0 by a pair of gradient field coils 1A and 1B. Coils 1A and 1B comprise, e.g., a pair of Helmholtz coils. A magnetic field intensity distribution in which a magnetic field intensity differs in accordance with the displacement along the z axis is obtained by magnetic gradient Gz along the z axis. Specific nuclei resonate with respect to static magnetic field H0 at angular frequency .omega.0 given by the following relation: EQU .omega.0=.gamma.H0
Where .gamma. is a magnetogyric ratio, which is inherent to each type of nuclei.
Excitation rotating field H1 at a radio frequency of angular frequency .omega.0 (generally, angular frequency .omega.0 is a radio frequency (RF)) which can resonate only specific nuclei is applied to patient P from a probe head. Since excitation rotating field H1 is applied not continuously for a long period of time but is applied in the form of pulses, it is called an excitation pulse. The probe head has a pair of transmission coils 2A and 2B comprising saddle coils, as shown in FIG. 1. Excitation pulse H1 is applied to patient P through coils 2A and 2B.
When excitation pulse H1 is applied, an MR phenomenon occurs only in slice portion S (which is a planar portion but has a given thickness in practice) selectively determined by magnetic gradient Gz along the z axis. Slice portion S is perpendicular to the z axis, as shown in FIG. 1. The MR phenomenon is detected as an MR signal by the probe head. The probe head also comprises a pair of reception coils 3A and 3B comprising saddle coils, as shown in FIG. 1, and the MR signal is detected by these coils 3A and 3B.
The MR signal detected by coils 3A and 3B is either a free induction decay (FID) signal or a spin echo signal. Which signal the coils 3A and 3B detect depends on the MR excitation method (or excitation pulse sequence) employed. In order to obtain a slice image consisting of MR data, two-dimensional position data on the plane of slice portion S is necessary. For this reason, to add the two-dimensional position data to the MR signal, magnetic gradient Gxy having gradients in various directions on the x-y plane is applied to field H0 by coils or the like (not shown) after slice portion S is excited to cause the MR phenomenon. The thus detected MR signal is subjected to predetermined processing, e.g., two-dimensional Fourier transformation, thereby reconstructing a slice image consisting of the MR data in slice portion S.
The obtained MR data does not separately include specific nuclear (e.g., hydrogen nuclear) density .rho., longitudinal relaxation time T1, and transverse relaxation time T2 (i.e. parameters inherent to the tissue of the patient P) as the MR data, but rather as a combination thereof. When a predetermined arithmetic operation is performed using MR images which are located at an identical position, have the same thickness, and satisfy predetermined conditions among MR images directly obtained from the MR signal, then one can obtain pure longitudinal and transverse relaxation times T1 and T2 and specific nuclear density p or a parameter image as an arbitrary combination of these parameters.
For example, a case will be described wherein an MR image is obtained from a spin echo signal by a spin echo method.
An MR original image obtained by the spin echo method, i.e., a spin echo image, includes specific nuclear density .rho., longitudinal relaxation time T1, transverse relaxation time T2 as parameters inherent to tissue, and repetition time TR (repetition cycle of a series of MR excitation - echo collection sequence) and echo time TE (an interval from when an excitation pulse is applied until a spin echo is observed) as parameters of imaging conditions, i.e., collecting conditions of MR (spin echo) signals. Specific nuclear density .rho., longitudinal relaxation time T1, and transverse relaxation time T2 as the parameters inherent to tissue are constant regardless of the imaging conditions, and repetition time TR and echo time TE as the parameters of the imaging conditions vary in accordance with the imaging conditions. For this reason, when an arithmetic operation is made using a plurality of spin echo images having different repetition times TR and echo times TE, parameter images for specific nuclear density .rho., longitudinal relaxation time T1, and transverse relaxation time T2 can be respectively obtained.
For example, as shown in FIG. 2, two spin echo images SE1 and SE2 are obtained when repetition time TR=2,000 msec and echo time TE=40 msec and of repetition time TR=2,000 msec and echo time TE=80 msec. Thus, parameter image PT of transverse relaxation time T2 can be calculated using these images SE1 and SE2.
When parameter image PT of transverse relaxation time T2 is formed by calculation using images SE1 and SE2, if calculation is simply made for each pixel, the value of a pixel which must have "0" is sometimes not "0" but an abnormal value. This is caused by background noise or an artifact of the image data during spin echo collecting process and/or spin echo image formation process.
In the conventional apparatus, as a countermeasure against this problem, a threshold value is set for data values of images, i.e., values for each pixel (gray levels), and pixels having values lower than the threshold value are set to have value "0".
With this method, however, since the threshold value is set based on experience, collected image data has a poor S/N. In particular, when the image data value of a background portion is large, the abnormal value of the background portion remains without being suppressed. In contrast to this, when the image data value is small, necessary image data may often be erased.